1. Field of the Invention
This invention relates to a method for improving imaging of a source of radiation and to a device for imaging a source of radiation, and more specifically, this invention relates to a method and device for producing a high spatial resolution three-dimensional image of a source of x-ray and gamma-ray radiation for medical and other application by using a plurality of diffracting crystals, collimators, and detection devices.
2. Background of the Invention
Cancer tumor cells have high rates of metabolism and multiply rapidly. Substances injected into the body tend to migrate to locations of such high growth and become incorporated in this new growth. If the injected substance incorporates a short-lived radioactive isotope, the location of a tumor can be detected by locating the region of high radioactivity. Aside from pinpointing tumor location, an image of the tumor is also desirable to ascertain its shape, size, and juxtaposition with adjacent structures. For many medical applications it is imperative that a tumor be detected as early as possible, and early tumors are very small in size. Thus their detection and identification requires the ability to image very small sources. Also, medical research often uses small animals, with very small organs, and the availability of devices with very high spatial resolution is of the utmost importance.
One method used to detect tumors is to first inject a body with a biological substance that contains radioactive compounds such as the Technetium isotope 99mTc, which is a 140.5 kiloelectronVolt (keV) gamma emitter having a half-life of 5.9 hours. The gamma rays are detected by a large sodium iodide (Nal) scintillator crystal placed behind a collimator grid yielding at best an 8 millimeter (mm) resolution at the location of the source. The scintillator is viewed by a plurality of photomultiplier tubes and the location of a scintillation event is determined by a computer analysis of the relative intensity of the photomultiplier signals. The collimator/scintillator assembly is placed above and very close to the patient. Aside from this method yielding a low resolution of between approximately 8 mm and 1 centimeter (cm), the image produced is limited to the plane parallel to the surface of the scintillator. As such, the technique provides no depth information about the source. This deficiency can be remedied somewhat by adding another collimator/scintillator assembly below the patient, comparing the counting rate of the two scintillators, and thus estimating the position of the source along the line joining them. In the latest revision of this method the large Nal detector plus collimator is rotated around the patient, taking a plurality of images at different angles. This allows one to generate a three-dimensional image of the radiation emitting area. There are considerable additional costs associated with this method and the fact that this method has been introduced in spite of the additional costs underscores the importance of three-dimensional imaging.
Another popular imaging technique is positron emission tomography (PET), used in diagnosis and medical research. In PET, a chemical compound containing a short-lived, positron-emitting radioisotope is injected into the body. The positrons (positively charged beta particles) are emitted as the isotope decays. These particles annihilate with electrons in surrounding tissue. Each annihilation simultaneously produces two 511 (keV) gamma rays traveling in opposite directions. After passing through collimators, these two gamma rays are detected simultaneously by scintillation detectors placed at 180 degrees to each other, and on opposite sides of the patient. The signals from the detectors' photomultiplier tubes are analyzed by a computer to facilitate the production of an image of the radiation-emitting region.
Numerous drawbacks exist with scintillation detector tomography. For instance, the typical coarse resolution of no less than 8 mm often results in smaller structures being overlooked. This prevents early detection of cancerous tumors when they are least likely to have metastasized and when treatment is most likely to succeed. This is especially a disadvantage in the detection of breast cancer tumors wherein the tumors often become virulent before growing to a detectable size. Presently, mammography uses x-rays to detect tissue calcification. The assumption is made that this calcification is due to dead cancer cells and that there is a live cancer tumor in the immediate vicinity. Often however, there is no live tumor where calcification has been detected. In fact, the calcification may not have been due to a tumor at all. Unfortunately then, positive mammography results often lead to unnecessary surgical operations.
Also, because poor spacial resolution often causes images of actual small tumors to be diffuse, variations in background radiation are often mistaken for actual tumors, leading to unnecessary surgical operations. This inadvertent incorporation of background radiation is an artifact of scintillation detector use wherein the detector must be large enough to cover a given area of the body. Aside from intercepting the radiation emanating from the source under observation, however, the large detectors also detect all ambient background radiation penetrating the scintillating region and this ambient radiation is analyzed as if it had been emitted by the source under observation.
Another drawback to using imaging techniques incorporating scintillation detectors is that all of the various radiations emitted by the source are detected by the detectors. As such, a specific radiation having an energy indicative of a specific, injected isotope cannot be easily scrutinized.
Recently, efforts have been made to improve scintillation detector tomography. Some PET instruments now achieve a resolution as small as 4 mm. Such improvements entail considerable expenditures and have the additional drawback that the improvement in resolution has come at the cost of a decrease in counting rate. This entails in turn either a longer examination time per patient or the injection of a stronger dose of radiation. Furthermore, the prospects for further improvements in resolution are limited by the fact that such improvements require collimators with ever smaller apertures, and therefore greater mass, together with lower count rates. This increase in collimator mass will increase the number of forward Compton-scattered photons in the collimators and these forward scattered photons are often indistinguishable from those emanating directly from the source.
Significant improvements in spatial resolution and in detection efficiency as well as a three dimensional location of the source using a crystal diffraction method for focusing the radiation emanating from the source was disclosed in U.S. Pat. No. 5,869,841 (1999) (granted to the same inventor as the present invention and assigned to the same assignee) and incorporated herein by reference. The enhanced focusing provided in the '841 patent allows the use of much smaller amounts of radioactive substances in order to locate features of interest in the patient. Experiments at the inventor's laboratory have demonstrated the effectiveness of this method and have achieved a spatial resolution of 7 mm. While this is adequate under many circumstances, better spatial resolution would provide significant advantages.
Thus a need exists in the art for an improved method and device for imaging x-ray and gamma-ray sources with sufficient spacial resolution to accurately observe structures smaller than 7 mm in size, even down to 1 mm in size or less. The invented method and the resulting device must have sufficient energy resolution to allow the imaging of radiation of a selected energy to the exclusion of others. The method and device also must limit the radiation to which the patient is exposed by incorporating a redirecting mechanism to detect radiation emanating from a tumor while disregarding ambient levels of radiation.